Cancer therapy via irradiation with energetic beams of protons or heavier ions has intrinsic physical advantages over the more widespread radiotherapy modalities based on X-rays. The ion beams deposit their energy much more selectively, primarily within a small volume (referred to as the Bragg peak) near the endpoint of the beam particle's path within the patient. This permits high radiation doses to be delivered to a tumor, while minimizing the dose to surrounding healthy tissue, and especially to sensitive organs that may be adjacent to the tumor volume. This physical advantage greatly reduces long- and short-term toxicity side effects of the radiation treatment.
In order to take full advantage of the ion beam benefits, a new generation of ion beam therapy technology is being developed, to make the treatment faster, more precise, simpler, cheaper and requiring much smaller footprints, so that it becomes economically more competitive with X-ray therapy. These new ion beam therapy facilities require advanced technical developments in the accelerators that produce the beams, in the gantry systems that deliver the beams to a patient from multiple angles, in the system that scans a beam of small cross-sectional area (a so-called “pencil beam”), and of adjustable energy and intensity, over the tumor volume and in the detector systems that carefully monitor and provide feedback on the delivered dose.
A central aspect of these new facilities is Intensity Modulation, typically provided by the use of so-called “spot beam scanning” (SBS). In this approach, an intense beam of small (sub-centimeter) diameter is used to deposit energy in a very short (no more than several milliseconds) time interval, primarily in a very small volume (“voxel”) near the back (distal) edge of a tumor. Electromagnets are then used to move the beam transversely in 2 dimensions to irradiate a neighboring voxel. With many more such steps, the entire lateral dimension of the deepest part of the tumor is irradiated. Then the beam energy is reduced and a similar transverse scan, though generally of different lateral extent, deposits energy primarily in a tumor slice slightly less deep within the patient. This sequence is repeated until the entire tumorous volume is “painted” with radiation. The beam intensity can be varied with each voxel in the entire scanning sequence to arrange for a radiation dose that conforms as closely as possible to the (arbitrary-shaped) tumor volume.
The SBS approach is more cost-effective and potentially more precise than conventional passive scattering treatments, with which it is compared schematically in FIGS. 1 and 2. But SBS is also more sensitive to organ motion. The interplay between the scanned beam frequency and the target motion frequency can result in localized under-dosage in parts of the target volume and over-dosage in other parts, an effect that has been seen in both simulations and experimental data. This sensitivity can be ameliorated in an alternative to the “pointillist” SBS approach, by scanning the ion beam continuously at a rate of at least a few cm/millisecond over a raster pattern in two dimensions, while varying the beam intensity continuously during the scan. The latter Raster Beam Scanning (RBS) approach can decrease the painting time of the tumor within a given depth layer by an order of magnitude compared to conventional SBS, so that multiple repaints of the depth layer can be performed within a time interval much shorter than typical organ motion periods in the patient's body. In either intensity modulation approach, SBS or RBS, the angle of beam entry can be changed by rotating the gantry for subsequent scans, to avoid radiation to critical organs and to spread any undesired radiation of healthy tissue outside the tumor over a larger volume.
Clinical application of pencil beam scanning in either of the above approaches requires accurate (to the 2% level) monitoring of the radiation dose delivered to each voxel, with the possibility of rapid feedback to the control system to alter the irradiation plan for subsequent voxels or subsequent repaints or subsequent tumor layers. The ionization chambers in current widespread usage as ion beam dosimeters are not well suited to this purpose. They have shortcomings associated with non-linear response at the high beam fluxes (number of incident particles per second per square centimeter) foreseen, they do not provide dose measurements on the needed sub-millisecond time scales, and when they are sensitive to beam position, their measurements typically provide marginal spatial resolution for determining definitively that the beam has indeed moved from voxel to voxel or for pinpointing locations where beam may have been off during a raster scan.
The non-linearity is especially problematic for next-generation ion beam therapy facilities that will utilize superconducting synchrocyclotrons or alternative accelerators that deliver pulsed beams. These accelerators deliver the beam for each voxel in beam pulses that last typically no longer than ten microseconds, while succeeding pulses are separated by about 1 millisecond, thus amplifying the instantaneous beam flux to which the dosimeter is exposed by a factor of 100 or more. The non-linearity is further exacerbated for heavier ion beams, which yield higher ionization density within the ionization chambers. Furthermore, the relatively slow response time and marginal spatial resolution provided by the ionization chambers limits their ability to provide rapid feedback to the control system when hardware or software problems may produce unexpected dose delivery to a given voxel, or may interrupt a treatment that needs to be continued from the last irradiated voxel after the system is successfully restored.
Alternative dosimetry systems currently employed in ion beam therapy have similar limitations. Integrating detectors such as radiochromic films and alanine detectors have limited applications since they are not real-time devices. Scintillator screens viewed by CCD cameras represent the most popular type of real-time quality assurance (QA) detectors, typified by the Lynx PT detector commercially available from vendor, IBA. Although the scintillating screen provides excellent spatial resolution to identify beam position, its fundamental drawback is that the detector response is not linear with dose when the beam stopping power approaches its maximum value.
In radiochromic films, this non-linearity leads to underestimation of dose of up to 20% at the peak of the Bragg curve. In scintillator screens, as much as 20% quenching of the light output in a proton beam, 30% in an α-particle beam and 43% in a carbon beam has been observed at high ionization densities (i.e. at the Bragg peak). Furthermore, for on-line monitoring applications with scanning beams, the limited processing speed of modern CCD cameras (<40 frames per second) is too slow to match the requirements of sub-millisecond feedback. Gas Electron Multiplier (GEM) chambers with electronic readout are often used in nuclear and particle physics experiments to provide fast detectors with excellent spatial resolution, and hence, they have promise for in-beam dosimetry. However, they are known to provide output signal amplitudes that vary significantly with position, dose rate and dose history, making them inappropriate for absolute dose determination. Scintillating GEM detectors with optical readout have been proposed for quality assurance on scanning ion beam treatment plans, but again with CCD readout making them too slow for real-time feedback to the beam scanning control system.
Intensity-modulated ion beam therapy, delivered via either spot or raster beam scanning, thus requires cost-effective dosimeters that are as accurate as ionization chambers at moderate dose rates, while providing improved linearity at high dose rates, faster response time, and higher spatial resolution, still with an overall signal strength that varies negligibly when a fixed-intensity beam spot moves across the detectors.
Embodiments of the present invention provide such a single detector with the capability to satisfy all of these above-referenced requirements simultaneously, representing a much more cost-effective approach than would be obtained by combining multiple different detectors of the above-mentioned types, to overcome their individual limitations. These and other advantages of the invention, as well as additional inventive features, will be apparent from the description of the invention provided herein.